This invention relates generally to magnetic resonance imaging (MRI) using nuclear magnetic resonance (NMR) phenomena. It is particularly directed to a method and corresponding apparatus for more efficiently capturing and providing MR data for use in multi-dimensional imaging processes.
MRI is a widely accepted, medically important and commercially viable technique for obtaining digitized video images representative of internal body tissue and structures. There are many commercially available systems and there have been numerous publications describing their operation and other approaches to MRI. Many of these use multi-dimensional Fourier transformation techniques which are now well-known to those skilled in this art.
In general, MRI devices establish a constant homogeneous magnetic field together with a specific additional bias field gradient in a known plane or region under consideration to orient nuclear spins, and apply a radiofrequency pulse or a sequence of pulses to further condition spins or perturb the nuclei. Those nuclei in a known region of the bias field gradient emit an RF signal in a specific band determined by the magnetic field distribution, and these RF emissions are detected by receiving coils and the received signals are stored as a line of information in a data matrix known as the k-space matrix. The full matrix is built up by successive cycles of conditioning the spins, perturbing them, and collecting RF emissions. An image is then generated from this matrix by Fourier transformation, which converts the frequency information present in the RF oscillations to spatial information representing the distribution of nuclear spins in the tissue.
Magnetic resonance imaging has proven to be a valuable clinical diagnostic tool in a wide range of organ systems and pathophysiologic processes. Both anatomic and functional information can be gleaned from the MR data, and new applications continue to develop with each improvement in basic imaging technique and technology. As technologic advances have improved achievable spatial resolution, for example, increasingly fine anatomic details have been amenable to MR imaging and evaluation. At the same time, fast imaging sequences have reduced imaging times to such an extent that many moving structures can now be visualized without significant motion artifacts.
Often, however, there is a tradeoff between spatial resolution and imaging time, since higher resolution images require a longer acquisition time. This balance between spatial and temporal resolution is particularly important in cardiac MR, where fine details of coronary artery anatomy, for example, must be discerned on the surface of a rapidly beating heart. A high-resolution image acquired over a large fraction of the cardiac cycle will be blurred and distorted by bulk cardiac motion, whereas a very fast image acquired in a shorter time may not have the resolution necessary to trace the course and patency of coronary arteries. Some of the fastest imaging sequences currently implemented, such as echo planar imaging (EPI), approach the goal of yielding images of reasonable resolution in a suitably short fraction of the cardiac cycle. Other approaches have also been tried to eliminate the effects of cardiac motion, including k-space segmentation, in which image acquisition is divided up over several cardiac cycles with ECG gating to ensure that the heart is in the same phase of systole or diastole during acquisition of each segment. Cine images of multiple cardiac phases may be pieced together with this technique, with partial acquisitions of the signal data for different phases occurring in each cardiac cycle. One problem with this class of techniques is that respiratory motion can change the position of the heart over the course of several cardiac cycles. Partial acquisitions will then be misregistered, and artifacts will result. In an attempt to eliminate or adjust for respiratory motion, breath holds, respiratory gating, and navigator echo gating techniques have all been tried, and each of these techniques has had some significant successes. Nevertheless, an imaging strategy which allowed high-resolution images to be acquired comfortably within one or two phases of the cardiac cycle would circumvent many of the difficulties and residual artifacts associated with these compensation techniques.
The speed with which magnetic resonance (MR) images may be acquired has already increased dramatically over the past decade. The improvements in speed may be traced to a combination of advances in the technologies of magnet construction and actuation, and innovations in imaging strategy. Strong, fast-switching magnetic field gradients and fast electronics have allowed the intervals between data collections to be reduced significantly. Meanwhile, fast gradient-echo and spin-echo sequences have reduced image acquisition time by allowing greater portions of k-space to be sampled quickly. Echo planar imaging (EPI), fast low-angle shot (FLASH), turbo spin echo (TSE), and spiral imaging techniques all allow very short intervals between acquisition of successive data points. The DUFIS, OUFIS, RUFIS, and BURST family of sequences further reduce image acquisition time by eliminating time delays incurred during gradient switching and echo formation. Details of the above-mentioned eight techniques may be found in the following papers: P. Mansfield, Multi-planar image formation using NMR spin echoes. J Phys. C. 10, L55-58 (1977); A. Haase, J. Frahm,
D. Mattaei, W. Hanicke, K. D. Merboldt, FLASH imaging: rapid NMR imaging using low flip-angle pulses. J. Magn. Reson. 67, 256-266 (1986); J. L. Listerud, S. Einstein, E. Outwater, H. Y. Kressel, First principles of fast spin echo. Magn. Reson. Q. 8, 199-244 (1992); C. Meyer, B. Hu, D. Nishimura, A. Macovski, Fast spiral coronary artery imaging. Magn. Reson. Med. 28, 202-213 (1992); I. J. Lowe, R. E. Wysong, DANTE ultrafast imaging sequence (DUFIS). J Magn. Reson. Ser. B 101, 106-109 (1993); L. Zha, I. J. Lowe, Optimized ultra-fast imaging sequence (OUFIS). Magn. Reson. Med 33, 377-395 (1995); D. P. Madio, I. J. Lowe, Ultra-fast imaging using low flip angles and FIDs. Magn. Reson. Med 34, 525-529 (1995); and J. Hennig, M. Hodapp, Burst imaging. MAGMA 1, 39-48 (1993).
Increasing the speed of MR imaging further is a challenging proposition, since the aforementioned fast imaging techniques have already achieved an impressive efficiency. All these techniques allow very short intervals between acquisition of successive data points, and hence do not waste much time in accumulating the data to fill the k-space matrix of a size required to generate a given image. In flow-encoded EPI images, for example, the entire complex k-space matrix is filled in a single spin excitation (which is followed by multiple spin conditioning cycles involving the application of multiple stepped field gradients), and the resulting image matrix is likewise "full," with useful information stored in both the real and the imaginary channels. One common feature of nearly all the fast imaging techniques currently in common use, however, is that they acquire data in a sequential fashion. Whether the required data set, i.e., the k-space data matrix, is filled in a rectangular raster pattern, a spiral pattern, a rapid series of line scans, or some other novel order, it is acquired one point and one line at a time.
That is, the above-mentioned fast MR imaging has concentrated on increasing the speed of sequential acquisition by reducing the time intervals between scanned lines. Still, however, only a portion of k-space is acquired at a time, which sets a methodological upper limit to the achievable speed for data acquisition. Modifications to pulse sequences or to magnetic field gradients have yielded a gradual improvement in imaging speed by allowing faster sequential scanning of k-space, but these improvements face limits due to the length and number of the intervals necessary to create, switch or measure the magnetic fields or signals involved in data acquisition. It would therefore appear difficult to devise a purely sequential technique with significantly better efficiency than the current fast imaging techniques.
Parallel acquisitions, in which multiple lines of data are acquired simultaneously, offer a means of relaxing these constraints on imaging speed. A few proposals for parallel or partially parallel acquisition (PPA) of the magnetic response signals in MRI have been described in the literature. These techniques provide some reduction of MR scan time based on spatial encoding with multiple spatially distinct receiver coils. Multiple coil arrays have been used in MR imaging as reported in R. B. Roemer, W. A. Edelstein, C. E. Hayes, S. P. Souza, and O. M. Mueller, The NMR phased array. Magn. Reson. Med. 16, 192-225 (1990); C. E. Hayes and P. B. Roemer, Noise correlations in data simultaneously acquired from multiple surface coil arrays. Magn. Reson. Med 16, 181-191 (1990); C. E. Hayes N. Hattes, and P. B. Roemer, Volume imaging with MR phased arrays. Magn. Reson. Med. 18, 309-319 (1991). The increased information content of the multiple received signals in such systems has been used to increase the signal-to-noise ratio (SNR) of MR images. In PPA imaging, however, each array coil is characterized by a unique spatial response, so that each receiver adds spatial information to the localization process, and this information is used to reduce the number of spatial encoding gradient steps. For their successful operation, all of the PPA techniques rely upon accurate knowledge, or upon estimation, of the relative RF sensitivities of the component coils in the array used for imaging.
Among the PPA imaging schemes proposed to date using simultaneous data acquisition in multiple RF receiving coils, are those described in: J. W. Carlson, T. Minemura, Imaging time reduction through multiple receiver coil data acquisition and image reconstruction. Magn Reson Med 29, 681-688 (1993) and U.S. Pat. No. 4,857,846; J. B. Ra, C. Y. Rim, Fast imaging using subencoding data sets from multiple detectors. Magn Reson Med 30, 142-145 (1993); and Sodickson D. K., Manning W. J. Simultaneous Acquisition of Spatial Harmonics (SMASH): Fast Imaging with Radiofrequency Coil Arrays. Magn. Reson. Med 38: 591-603 (1997); Sodickson D. K., Bankson J. A., Griswold, M. A., Wright S. M. Eightfold improvements in MR Imaging speed using SMASH with a multiplexed eight-element array. Proc. Sixth Scientific Meeting of the International Society for Magnetic Resonance in Medicine, 577 (1998). These approaches have offered the promise of significant savings in image acquisition times.
Carlson and Minemura describe a two-fold acquisition time savings using two nested body coils. In their approach, partial data sets are collected simultaneously in the two coils, one of homogeneous sensitivity and the other with a linear gradient in sensitivity. Missing lines in k-space are generated using a series expansion in terms of other phase-encoded lines. This approach using body coils appears to require that a significant portion of the data for the partial k-space matrix be acquired before any of the missing lines can be filled in by postprocessing, and thus does not allow for the missing lines to be built up as the data arrives, in real time. The approach uses coil sensitivity information in place of some portion of the gradient phase encoding steps, but has drawbacks. The coils used by Carlson and Minemura are body coils, which provide large volume coverage but lower overall sensitivity than surface coils, and it would be difficult to augment their number to improve time savings.
The approach of Ra and Rim involves a simultaneous acquisition technique in which images of reduced FOV are acquired in multiple coils of an array and the Nyquist aliasing in those images is undone by reference to component coil sensitivity information. The unaliasing procedure involves a pixel-by-pixel matrix inversion to regenerate the full FOV from multiple copies of the aliased image data. The "subencoding" technique of Ra and Rim relies on estimates of component coil sensitivities by effectively probing the sensitivity at each pixel. This pixel-by-pixel approach can lead to local artifacts; for example, the matrix inversion can begin to fail in regions of low sensitivity. Further, by its very nature as a pixel by pixel dealiasing approach, the Ra & Rim method is computation-intensive and is limited to postprocessing, as it appears to require all image data to be present before the reconstruction can be undertaken.
The Sodickson et al technique, denoted by the acronym SMASH, which stands for SiMultaneous Acquisition of Spatial Harmonics was initially developed by one of the applicants herein, and is more fully described in U.S. patent application Ser. No. 08/446,358 filed Nov. 12, 1996, and in the above-referenced two Sodickson D. K., et al. papers; in Sodickson, D. K., Griswold M. A., Jakob P. M., Edelman R. R., Manning, W. J. Signal-to-noise ratio and signal-to-noise efficiency in SMASH Imaging. Proc. Sixth Scientific Meeting of the International Society for Magnetic Resonance in Medicine, 1957 (1998); and in further Abstracts by the same authors. The SMASH technique is a PPA technique which extracts additional spatial information from a collection of signals acquired simultaneously in multiple coils that have different sensitivities, by combining acquired signals with weights to synthesize multiple distinct signals each corresponding to a sinusoidal spatial variation in coil sensitivity. These spatial variations, or "spatial harmonics," take the place of spatial modulations normally produced by magnetic field gradients in a conventional MR imaging protocol, and the simultaneous acquisition of signals to form multiple spatial harmonics in the signal domain thus allows the simultaneous acquisition of multiple lines of MR data. The SMASH technique can be integrated with many of the fastest existing imaging sequences, yielding multiplicative improvements in imaging speed, and a factor of two to three time savings has already been demonstrated in vivo using SMASH with commercially available coil arrays, with up to eight-fold improvements reported in phantoms using specialized RF hardware. In principle, there is no limit to the number of k-space lines that may be scanned simultaneously, assuming that coil arrays with sufficient numbers of independent coil components having suitable sensitivities are available for a given field-of-view (FOV).
Although PPA imaging techniques based on coil arrays can provide a considerable improvement in imaging speed, the common constraint of the PPA imaging techniques is their dependence on the accurate measurement or knowledge of component coil sensitivities or sensitivity-related information. The PPA reconstructions rely upon an accurate estimate of the individual coil sensitivity functions in the underlying coil array. A number of strategies for coil sensitivity calibration have been proposed. Ra J. B., et al. Magn. Reson. Med 30: 142-145 (1993); Sodickson D. K., et al. Magn. Reson. Med. 38: 591-603 (1997); Axel L., et al. AJR 148: 418-420 (1987); Murakami J. W., et al. Magn. Reson. Med. 35: 585-590 (1996). These strategies basically fall into four groups.
First, the coil sensitivity profiles can be calculated from the Biot-Savart law using knowledge of the coil's size, shape and position relative to the slice-of-interest (SOI). However, this approach is generally impractical in vivo, since the theoretical field map may correlate poorly with the actual sensitivity profiles because of unpredictable variations in coil loading effects and inaccurate coil positioning.
Second, the sensitivity information can be obtained from images of a uniform phantom (i.e., a test object) taken at the same position as the in vivo images. However, the use of an in vitro reference can be problematic in many cases, since coil loading and/or coil position may change significantly from subject to subject, thereby changing the effective coil sensitivities existing in practice. In addition, acquiring and using these reference data to correct subsequent in vivo images can be impractical with arrays such as flexible phased arrays, since the exact locations of the individual coils may be affected by the subject's anatomy.
Third, an estimate of the coil sensitivity functions can be obtained by acquiring the required coil references in vivo in the desired image plane. This method was used for the first in vivo implementations of SMASH. However, this approach requires that a reference image set be acquired using an appropriate imaging technique, which we term "coil sensitivity weighted", before the postprocessing of the PPA images is possible. This procedure can be imperfect, since it requires first identifying a region of uniform spin density against which to calibrate the sensitivities. This requirement is often impossible to fulfill in vivo, especially in regions of highly varying spin density such as in the chest, which has regions of very low signal-to-noise in the area of the lungs, and regions of very high signal to noise in the area of the heart. In addition, B.sub.0 and B.sub.1 magnetic field inhomogeneities may distort the true coil sensitivity profiles, depending on the imaging technique used. Furthermore this procedure can also be time consuming, since it has to be performed for each SOI.
A fourth technique which avoids some of the problems mentioned above, is to derive an estimate of the sensitivity profiles from a combination of body coil and array coil images of the subject as reported by Murakami et al for phased array coils. In this approach, the surface coil image is divided by the body coil image to derive the array coil sensitivity profile. This approach accurately estimates the coil sensitivity functions in areas where sufficient signal-to-noise ratio is available for the quotient to be meaningful. However this condition may fail to hold in areas such as the lungs. This method also increases scan time significantly, since the data has to be acquired in additional independent scans. In addition, this approach is difficult to apply in moving tissue structures, since the body coil and phased array coil image have to be obtained ideally in exactly the same position. Therefore, in the situation of involuntary subject motion, breathing and cardiac motion, the accuracy of this coil sensitivity calibration can be impaired or entirely defeated. Ra and Rim have described a similar method using a reference array image set without the body coil image, but in other respects it suffers from the same difficulties.
In summary, PPA techniques, including SMASH, rely upon accurate estimation of the sensitivity functions of individual coils in a coil array. Estimation can be a cumbersome, inaccurate and time-consuming procedure which in the worst case may eliminate the time savings which PPA techniques promise. The problems posed by the need to determine coil sensitivities limits potential applications of faster imaging with PPA.
It is therefore desirable to develop improved methods and devices for determining relevant parameters for image reconstruction without cumbersome determinations of exact receiving coil sensitivity.
It is also desirable to provide a method and apparatus for determining such parameters or information from in vivo signals, and from such signals acquired with little or no additional spin conditioning.
In order to address these limitations, applicants have developed a new internal calibration technique for SMASH imaging, called AUTO-SMASH, in which sensitivity-related reconstruction parameters are determined from signals aquired (e.g., during the actual scan) by an auto-calibration mechanism that relies on fitting sets of collected signals. Details of both acquisition and reconstruction strategies in this new AUTO-SMASH approach are provided below, together with a description of aspects of the invention relating to extraction of calibration information for SMASH and other PPA imaging techniques.